The invention relates to gamma-ray camera systems, in particular to spectral processing of data in gamma-ray camera systems.
FIG. 1 schematically shows in vertical cross-section a gamma-ray camera system 2 viewing a sample 4. In this case the sample comprises a point-like gamma-ray source 6 embedded within an extended body 8. The gamma-ray camera system comprises a gamma-ray camera 10 and an energy spectra accumulating component 12. The gamma-ray camera 10 is an Anger-type camera [1]. This is a type widely used as a diagnostic tool in nuclear medicine. The gamma-ray camera 10 includes a gamma-ray imager 14 coupled to a detector read-out component 16. The energy spectra accumulating component includes an energy spectra accumulator 18 and a data storage component 20.
The gamma-ray imager 14 includes a parallel collimator 22, a scintillator crystal 24, a light guide element 26 and a plurality of photo-multiplier tubes 28. The scintillator crystal is, for example, a large single crystal of Thallium doped Sodium Iodide (NaI(Tl)). The scintillator crystal is shielded from gamma-ray photons not incident through the parallel collimator by a shield 30.
The gamma-ray imager provides a 50 cm square image plane. The parallel collimator comprises an array of apertures with a characteristic cell size of 2 mm. Each cell provides a field-of-view of around 3°. The photo-multiplier tubes form a close packed hexagonal array of 61 tubes arranged to collectively view much of the 50 cm square scintillator crystal 24 forming the image plane.
During an exposure period, which in nuclear medicine imaging applications might typically be around 5 minutes, photons are emitted by the gamma-ray source in all directions. In a typical application, the photons will be emitted by radio-labelled pharmaceuticals in a patient's body (i.e. pharmaceuticals labelled with a radioactive tracer). In the example shown in FIG. 1, the gamma-ray source is Cobalt-57. Cobalt-57 primarily emits gamma-ray photons with an energy of around 122 keV. Six such photons, labelled A-F, are emitted in the plane of the figure as schematically shown in FIG. 1. Photons A, B and C exit the sample 4 in the directions indicated in the figure and are not seen by the gamma-ray camera. Photon D is emitted towards the gamma-ray camera, but is not sufficiently parallel to the axis of the parallel collimator 22 to pass through it. As can be seen from the figure, photon D is absorbed in a wall of the parallel collimator, and as such is not detected by the scintillator crystal 24. Photon E, however, does reach the scintillator crystal 24 since its path is within the parallel collimator's field-of-view. The energy of photon E is deposited in the scintillator crystal 24 in a scintillation event. A detection of this kind, where the gamma-ray photon travels directly between the gamma-ray source and the scintillator crystal, is known as a direct detection event. The detection event generates a pulse of optical radiation which illuminates several of the photo-multiplier tubes 28 via the light guide element 26. The light guide element assists in coupling the pulse of optical radiation from the scintillator crystal, which typically has a relatively high refractive index at visible wavelengths. In a typical scintillation event, the resulting pulse of optical radiation will be detected by up to seven of the photo-multiplier tubes.
The signals from the photo-multiplier tubes are supplied to the detector read-out component. The detector read-out component determines the X- and Y-coordinates of the scintillation event from the relative intensities of the signals seen by each of the photo-multiplier tubes. The detector read-out component also calculates the total energy deposited in the scintillation event from a summation of the signal amplitudes seen by the photo-multiplier tubes. Read-out components for Anger-type gamma-ray cameras are well known [1]. One mode of operation is known as list-mode operation. In this mode the detector read-out component outputs a signal in response to each scintillation event, the output signal including the calculated energy deposited in the scintillator crystal and the X- and Y-coordinates of the scintillation event. The output signals from the detector read-out component are coupled to the energy spectra accumulating component 12.
The functionality of the energy spectra accumulating component 12 in this example is provided by a suitably programmed general purpose computer. The computer includes an appropriate interface to receive and decode the output signals from the detector read-out component 16. The energy spectra accumulator 18 within the energy spectra accumulating component 12 operates to generate a three-dimensional observed data array I(X, Y, E). This array comprises a count of the number of scintillation events occurring within an exposure period as a function of their X, Y position within the detector, and the energy deposited in the scintillator crystal. A bin-size used in generating I(X, Y, E) is typically 5 mm for each spatial coordinates (i.e. X and Y) and 2 keV for the energy coordinate (i.e. E). Although the scintillator crystal 24 is a large single crystal, the 5 mm spatial binning used in generating I(X, Y, E) defines effective detector pixels, and these have a size of around 5 mm square. However, because of Compton scattering within the sample 4, the resolution in a resulting image is worse than this.
When an exposure is completed, the observed data array I(X, Y, E) is normalized to the exposure time and stored in the data storage component 20 of the energy spectra accumulating component for later analysis. I(X, Y, E) will typically be used in generating two-dimensional diagnostic images representing gamma-ray emission intensity from the source 4 as seen within selected gamma-ray energy ranges.
FIG. 2 shows a typical energy spectrum which might be seen in one of the detector pixels of a gamma-ray camera system similar to that shown in FIG. 1. Count rate n(E) is plotted as a function of energy E for a detector pixel at position X0, Y0—i.e. n(E)=I(X0, Y0, E). In this example, an isolated (i.e. not embedded in a body) gamma-ray point source is viewed to show the intrinsic energy resolution of the gamma-ray camera. The gamma-ray source is again Cobalt-57 emitting primarily at a photo-peak energy of E0=122 keV. In FIG. 2 the full width at half maximum (FWHMINT) of the peak corresponding to gamma-ray emission from the Cobalt-57 source is approximately 25 keV. Accordingly, at an energy of 122 keV the gamma-ray camera has an intrinsic energy resolution of around 20%. This relatively poor energy resolution is significantly worse than that predicted by photon-statistics alone and is due to several factors. One effect is the variance in the scintillation efficiency of the crystal itself, this is energy dependent and cannot be corrected for simply. Another effect is the non-uniformity of the response of the photo-multiplier tubes. This means scintillation events which deposit the same energy in the scintillator crystal will appear to have different energies depending on which photo-multiplier tubes are illuminated, and how the actual illumination falling on each individual photo-multiplier tube is distributed. A further effect is the variance in the light-collection efficiency of the scintillator crystal and photo-multiplier tube assembly for events which occur at different locations within the detector crystal.
When forming a diagnostic image from the stored data array I(X, Y, E), for example in order to represent the distribution of Cobalt-57 within the field-of-view of the gamma-ray camera, the relatively poor intrinsic resolution is not necessarily a significant problem with an isolated point-like gamma-ray source. This is because for each pixel in the image, the intensity of gamma-ray emission seen by the gamma-ray camera can be represented by a summation over the full width of the peak around 122 keV, for example between 100 keV and 150 keV. Because the poor energy resolution does not directly effect the imaging capabilities of the camera in such cases, a summation over this wide energy range provides the best signal-to-noise ratio possible by making use of all detected events, without unduly compromising image quality.
In practice, however, it will generally be necessary to generate diagnostic images of gamma-ray emission within the field-of-view of the gamma-ray camera by summing the energy spectra recorded in each pixel over a more restricted energy range than the full width of the peak seen in FIG. 2. This can be important, for example, to improve the signal-to-noise ratio in a derived diagnostic image, or to distinguish between different gamma-ray emission energies which can be used in some clinical observations. For example, in clinical studies Technetium-99 (which emits primarily at 140 keV) and Thallium-201 (which emits primarily at 80 keV) are sometimes used as radioactive tracers to examine the relative distribution of different pharmaceuticals in a patient. To differentiate spectrally between multiple gamma-ray sources in the field-of-view, it will not be possible to sum over all of the scintillation events relating to, say, the Thallium-201 emission without also including some of the scintillation events relating to the Technetium-99 emission due to the spectral overlap of the observed peaks. This means that it is not normally possible to make simultaneous observations of multiple radioactive tracers, and sequential observations must be made.
In cases where the field-of-view includes only a monochromatic source of gamma-ray emission, the poor energy resolution can still lead to degradation in derived diagnostic images where the source is embedded in a body. This is due to scattering, for instance Compton scattering, in the body surrounding the gamma-ray source.
The energy of gamma-rays emitted by radioactive sources used to label commonly-used pharmaceuticals in nuclear medicine is typically on the order of 102 keV. For example, Technetium-99 emitting at 140 keV is commonly used. This energy is chosen to be sufficiently energetic to allow emitted gamma-ray photons to escape from the surrounding body in which the gamma-ray source is embedded, yet without being so energetic as to make collimation and detection difficult. One disadvantage of this choice of energy is that the gamma-ray photons have a relatively high probability of scattering within the surrounding body before being viewed by the gamma-ray camera. An example of such a scattering event is shown by the photon labelled F in FIG. 1.
Referring to FIG. 1, photon F is initially emitted by the gamma-ray source 6 in a direction away from the gamma-ray camera system 2. Photon F should not normally contribute to the observed data array I(X, Y, E). However, in the case shown in FIG. 1, photon F undergoes a Compton scattering event in the surrounding body 8 which scatters it towards the gamma-ray camera. The scattered photon passes through the parallel collimator 22 and is detected by the scintillator crystal 24. Photon F is scattered off an electron marked e− in FIG. 1. The electron carries away some of the energy of Photon F. As a result of the Compton scattering of photon F, the gamma-ray camera system records a scintillation event occurring at a position not commensurate with the position of gamma-ray photons arriving directly from the source 6, and at an energy slightly lower than of the gamma-ray photons arriving directly from the source 6. While for simplicity a point source of gamma-ray photons is shown in FIG. 1, in general there will be an extended gamma-ray source within the surrounding body. This means that not only are gamma-ray photons from an individual portion of an extended source mapped to a particular example pixel scattered into other pixels, but gamma-ray photons from other portions of the extended gamma-ray source are scattered into the example pixel. As a consequence, an additional source of background noise in introduced throughout the image. The additional background noise not only leads to a reduction in the sensitivity of the camera due to the reduced signal-to-noise ratio in each pixel, but also impacts on the spatial resolution of derived images as a Compton scattered halo becomes associated with each portion of the gamma-ray source due to the non-direct detection events. Non-direct detection events are known as Compton scattered detection events.
In order to minimise the deleterious effect of Compton scattering in the source, it is necessary to reject as many of the Compton scattered detection events as possible, without unnecessarily discarding too many of the direct detection events. As noted above, when a gamma-ray photon is Compton scattered, some of its energy is imparted to a free electron in the scattering material. This means that, in principle, it is possible to distinguish between direct detection events and Compton scattered detection events on the basis of the energy deposited in the scintillator material comprising the scintillator detector. However, the poor energy resolution of scintillator crystal based gamma-ray camera systems is not able to provide suitable energy discrimination at 140 keV to make this an efficient process.
FIG. 3 shows a typical energy spectrum which would be seen in one of the detector pixels of a gamma-ray camera system similar to that shown in FIG. 1. As with FIG. 2, count rate n(E) is plotted as a function of energy E for a single pixel at position X0, Y0—i.e. n(E)=I(X0, Y0, E). The gamma-ray source is again Cobalt-57 emitting primarily at an energy of E0=122 keV. However, in the example shown in FIG. 3, data is obtained with the gamma-ray source positioned behind a 5 cm thick body of water, as opposed to in isolation. The body of water corresponds to the surrounding body 8 shown in FIG. 1, and it is in this body of water that Compton scattering can occur leading to Compton scattered detection events being detected.
On the high energy side of the peak seen in FIG. 3 the detected count rate profile is generally similar to that seen in FIG. 2. However, on the low energy side there is a significant increase in count rate compared to the same region of the spectrum shown in FIG. 2. The excess count rate in this region of the spectrum is due to Compton scattered detection events. In order to highlight the magnitude of the contribution of these events, an estimate of the profile which would be seen in the absence of Compton scattering is marked with a dashed line in FIG. 3. The corresponding area under the curve shown in FIG. 3 estimated to derive from only direct detection events is identified by hatching. The un-hatched area under the curve, which is marked CS, reflects the contribution to the count rate due to Compton scattered detection events. In generating a diagnostic image to represent the distribution of Cobalt-57 in the field-of-view from data such as shown in FIG. 3, it is necessary to determine an appropriate energy-width, known as an energy window, over which to sum the count rate n(E). The energy window must be chosen to obtain a summed count rate reflecting as many direct detection events as possible, while rejecting as much of the contribution from Compton scattered detection events as possible. Determining the most appropriate width of energy window will generally be a matter of compromise. For example, if an energy window such as that marked W1 in FIG. 3 were to be used, most of the direct detection events would be included in the image generation as desired. However, with this wide energy window a significant fraction of Compton scattered detection events will also be included. If, on the other hand, a narrower energy window were to be employed, such as the one marked W2 in FIG. 3, there would be a significant reduction in the number of Compton scattered detection events included in the summation, both in absolute and relative terms. However, with a narrower energy window there would also be a reduction in the number of direct detection events included. This leads to derived diagnostic images with a poor signal-to-noise ratio.
FIG. 4A schematically shows in negative an idealized diagnostic image of an example extended gamma-ray source distribution obtained using an idealized gamma-ray camera system. The example source distribution is in the form of a cross with a bright spot at the centre, and the source distribution is directly reflected in the resulting image.
FIGS. 4B and 4C schematically show how the same example gamma-ray source distribution imaged in FIG. 4A would appear when imaged with a gamma-ray camera system of the kind discussed and using different width energy windows.
FIG. 4B shows the result of using a wide energy window, such as the one marked W1 in FIG. 3. As outlined above, the inclusion of a significant fraction of Compton scattered detection events leads to relatively poor spatial resolution in the generated image due to the Compton scattered halo appearing to surround each element of the gamma-ray source. FIG. 4C shows the result of using a narrow energy window such as the one marked W2 in FIG. 3. As outlined above, the rejection of a significant fraction of non-Compton scattered events leads to a reduction in the appearance of the Compton scattered halo, but with a correspondingly reduced signal-to-noise ratio. This causes the image to appear faint against a relatively high background noise level.
As a consequence of the failings of gamma-ray camera systems of the type described above, alternative designs for gamma-ray camera systems which help to improve the rejection of Compton scattered detection events while maintaining the direct detection events have evolved.
For example, a number of research groups are seeking to exploit the improved energy resolution characteristics of room-temperature semiconductor detectors such as CdZnTe. These are capable of providing an intrinsic full width at half maximum energy resolution of around 4% at 140 keV [2]. This compares favourably with the approximately 20% seen above for an Anger-type gamma-ray camera system. The improved energy resolution allows a narrower energy window to be employed whilst retaining many more of the direct detection events since these are now contained in a narrower spectral peak. However, there are a number of disadvantages associated with this type of design. For example, each detector pixel requires a separate read-out electronics channel, the detection efficiency is relatively low and the camera costs are high.
Improved energy resolution has also been achieved in the recently introduced “2020tc Imager” camera manufactured by Digirad, San Diego, USA. This uses silicon positive-intrinsic-negative (PIN) diodes to record the scintillation flash generated in a large array of discrete Caesium Iodide (CsI) scintillation crystal elements. However, this design again requires significantly increased complexity in the read out electronics in comparison to other types of gamma-ray camera system, such as the kind discussed above.
Methods of improving the spectral resolution of non-imaging scintillator-based gamma-ray detectors by post-exposure spectral processing have also been described [3, 4, 5]. These techniques involve determining a detector response function for the single element detector and deconvolving this response function from observed spectra. However, these techniques have been applied only to single element specific scintillator geometries so as to provide a calculable transfer function describing the response of a photo-detector to pulses of optical photons from a scintillation event.
Various other schemes for reducing the effect of Compton scattered detection events in gamma-ray camera systems have been attempted.
U.S. Pat. No. 5,903,008 describes the use of dual energy-windows for diagnostic image formation [6]. This dual energy-window technique measures the relative number of counts recorded in two energy channels; one centred on the photo-peak energy and the second, some 10-15% below that energy. This ratio is measured initially when there is no scattering present. Thereafter, the photo-peak values are modified according to the value of the number of counts in the lower window. This value is dominated by the presence of scattered events.
U.S. Pat. No. 5,530,248 describes a scheme for theoretical modelling of the contribution of Compton scattered detection events to the energy spectra [7]. This technique generates a trial function from the energy spectrum recorded in each pixel when no scattering material is present. This is achieved by taking the first differential of the spectrum to emphasise the photo-peak. This modified function is then fitted to the energy spectra seen when a scattering medium is present in order to distinguish between ‘wanted’ and ‘unwanted’ photons.
U.S. Pat. No. 5,633,499 describes a scheme based on calculating a correction table to be applied to an image acquired using a conventional method of selecting only those events that fall within an energy window spanning ±10% of the photo-peak [8]. The correction value is derived from the measurement of the centroid of the spectrum recorded in a particular detector pixel. The data are used to estimate the scatter contribution.
U.S. Pat. No. 5,561,297 describes a spectrum subtraction method [9]. This method in essence uses the idea of subtracting a reference spectrum for each detector pixel form an observed spectrum. The subtracted reference spectrum is that recorded in the absence of a scatterer.